Implantable tissue perfusion sensing system and method

ABSTRACT

A medical device for sensing cardiac events that includes a plurality of light sources capable of emitting light at a plurality of wavelengths, and a detector to detect the emitted light. A processor generates an ambient light measurement in response to ambient light detected by the detector, generates a plurality of light measurements in response to the emitted light detected by the detector, and adjusts the plurality of light measurements in response to the ambient light measurement.

RELATED APPLICATION

The present application claims priority and other benefits from U.S.Provisional Patent Application Ser. No. 60/892,033, filed Feb. 28, 2007,entitled “IMPLANTABLE TISSUE PERFUSION SENSING SYSTEM AND METHOD”,incorporated herein by reference in its entirety.

CROSS-REFERENCE TO RELATED APPLICATION

Cross-reference is hereby made to the commonly-assigned related U.S.Applications, attorney docket number P31790.00, entitled “IMPLANTABLETISSUE PERFUSION SENSING SYSTEM AND METHOD”, to Cinbis et al., attorneydocket number P31791.00, entitled “IMPLANTABLE TISSUE PERFUSION SENSINGSYSTEM AND METHOD”, to Cinbis et al., attorney docket number P31792.00,entitled “IMPLANTABLE TISSUE PERFUSION SENSING SYSTEM AND METHOD”, toCinbis et al., and attorney docket number P31793.00, entitled“IMPLANTABLE TISSUE PERFUSION SENSING SYSTEM AND METHOD”, to Cinbis etal., incorporated herein by reference in their entireties.

TECHNICAL FIELD

The present invention relates generally to sensing cardiac signals in amedical device, and more particularly, the present invention relates toimplantable sensors for detecting tissue perfusion to verify detectionof a cardiac event in response to sensed cardiac signals.

BACKGROUND

Wide assortments of implantable medical devices are presently known andcommercially available. These implantable medical devices include avariety of implantable cardiac devices. For example, implantable pulsegenerators (IPGs) are a type of cardiac device that is generally used toelevate the heart rate that is beating too slowly. This type of deviceis sometimes referred to as a bradycardia device or a pacemaker. Anothertype of implantable cardiac device is an implantable cardiacdefibrillator (ICD). This type of device, often referred to as atachycardia device, is generally used to provide burst pacing pulses ora defibrillation shock to the heart when the heart is beating too fastor goes into fibrillation. Another type of device is a cardiacresynchronization device used to treat heart failure.

Each of these types of implantable cardiac devices includes a sensor orsensors to monitor the patient's cardiac system to facilitatedetermination of when and what action to take. For example, manyprevious implantable cardiac devices have relied upon electrical sensorsextended into the right ventricle of the heart. These electrical sensorsmeasure the electrocardiogram (ECG) signal in the heart to determine howwell the heart is functioning, and to determine what, if any, action theimplantable cardiac device needs to take. Unfortunately, extending alead into the heart, or attaching a lead to the outside of the heart isa relatively invasive procedure and is thus not desirable for allpatients. Without a sensor into the heart it can be problematic toeffectively monitor the patient's cardiovascular status. Specifically,the presence of spurious electrical signals caused by muscle movementand other factors can interfere with attempts at cardiac monitoringusing sensors that are not extended into or attached to the outside ofthe heart itself.

Thus, there remains a need for additional implantable sensing techniquesfor monitoring a patient's cardiovascular status.

BRIEF DESCRIPTION OF THE DRAWINGS

A more complete understanding of the invention may be derived byreferring to the detailed description and claims when considered inconjunction with the following figures, wherein:

FIG. 1 is a schematic diagram of electronic circuitry included in amedical device according to an embodiment of the invention;

FIG. 1A is a schematic view of a tissue perfusion sensor system inaccordance with a embodiment of the invention;

FIG. 2 is a side view a tissue perfusion sensor system in accordancewith a embodiment of the invention;

FIG. 3 is a side view a tissue perfusion sensor system in accordancewith a embodiment of the invention;

FIG. 4 is a flow diagram of a method of detecting a change in perfusionin accordance with one embodiment of the invention;

FIG. 4A is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention;

FIG. 4B is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention;

FIG. 4C is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention;

FIG. 4D is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention;

FIG. 4E is a flow chart of a method for correcting sensing by a medicaldevice for the effects of tissue encapsulation according to anembodiment of the invention;

FIG. 4F is a graphical representation of the optical pulse amplitude(AC) and the average optical amplitude (DC) associated with the methodfor correcting sensing by a medical device for the effects of tissueencapsulation of FIG. 4E according to one embodiment of the invention;

FIG. 4G is a flowchart of a method for correcting sensing by a medicaldevice for the effects of tissue encapsulation according to anembodiment of the invention;

FIG. 4H is a flowchart of a method for generating a blood volume indexin a medical device according to an embodiment of the invention;

FIG. 4I is a flowchart of a method for generating an oxygenation indexin a medical device according to an embodiment of the invention;

FIG. 5 is a graphical view of measured light signals in accordance withone embodiment of the invention;

FIG. 6 is a graphical view of compensated measurements in accordancewith one embodiment of the invention;

FIG. 7 is a graphical view of uncompensated IR time intervals inaccordance with one embodiment of the invention;

FIG. 8 is a graphical view of a first light measurement and a secondlight measurement taken during fibrillation and after defibrillation;and

FIG. 9 is a graphical view of the isobestic volume index and the redoxygenation index calculate according to an embodiment of the invention.

DETAILED DESCRIPTION

The following detailed description is merely exemplary in nature and isnot intended to limit the invention or the application and uses of theinvention. Furthermore, there is no intention to be bound by anyexpressed or implied theory presented in the preceding technical field,background, brief summary or the following detailed description.

The invention provides a sensor system and method for monitoring changesin tissue perfusion that is adaptable for use in medical devices,including implantable and external medical devices. The tissue perfusionsystem and method provides the ability to determine if nearby tissue isbeing adequately perfused. In general, perfusion is a function of bloodvolume, blood pressure, oxygen content and flow. Specifically, bydetermining if tissue oxygenation and/or blood volume is beingmaintained, the system and method can be used, along with other sensormeasurements, to determine what action, if any, the implantable medicaldevice should take. Additionally, the system and method can be adaptedto be tolerant of noise sources such as mechanical noise and tissueencapsulation.

In one embodiment, the tissue perfusion sensor system includes at leasttwo light sources and a light detector. The first light source provideslight at a wavelength where light absorption in the tissue is dependentupon the oxygen saturation level of the hemoglobin and myoglobin in thetissue as well as the total volume of arterial and venous blood in thetissue. The second light source provides light at a wavelength wherelight absorption in the tissue is substantially independent of theoxygen content in the blood, but where the light absorption is dependentupon the blood volume in the tissue. Light from the first and secondlight sources are emitted into the surrounding tissue and received backat the light detector after transmitting through, and/or being reflectedby the surrounding tissue. A sensor controller receives the lightmeasurements from the light detector and compensates the measurement ofdetected first light using the measurement of detected second light.

Specifically, a change in the measurement of first received lightcorresponds to a change in the overall oxygen content of the tissue aswell as change in the volume of the blood in the tissue. Therefore, thatchange can be the result of a change in oxygenation (i.e., a change inthe oxygen content of the tissue) or a change in the volume of bloodcaused by a change in arterial pressure, vasoconstriction or dilation, achange in posture, or muscle motion. A change in oxygenation or arterialpressure would generally indicate a change in the patient's cardiachealth due to a change in tissue perfusion. Due to the dependence of thereceived first light to muscle motion or vasoconstriction/vasodilation,it cannot, by itself, be used to reliably detect perfusion in thetissue. However, because the second light source was chosen to have awavelength where absorption in the tissue is independent of oxygencontent in the blood, but is dependent upon blood volume, it can be usedto separate the effects of muscle motion and blood volume change fromthe blood oxygen change in the received first light measurement. As oneexample, by scaling the second light measurement by a gain constant,then subtracting the scaled second measurement from the first lightmeasurement, the result is a measurement that will track a change intissue oxygenation substantially independent of blood volume changes.Thus, when using two light sources as described the sensor system andmethod can monitor oxygenation in tissue to determine if oxygen contentin the blood is being maintained. Alternatively, the two signals can beprocessed to remove the effects of the change in oxygen content andmotion artifacts, leaving a signal dependent on blood volume.

In another variation on this embodiment a third light source is providedto further improve the diagnostic ability of the sensor. Specifically, athird light source is added that provides a second distinct wavelengthof light dependent upon the oxygen content of blood in the tissue. Inparticular, the wavelengths are chosen such that one is independent ofthe oxygen content, one wavelength is shorter than the oxygenindependent wavelength, and the other wavelength is longer than theoxygen independent wavelength. When so selected, the tissue oxygenresponse from the third light source is expected to change in oppositedirection with respect to the response from the first light source (i.e.if the signal from the first light sensor is increasing due to a changein tissue oxygen, the signal from the third light sensor will bedecreasing). The sensor controller receives the light measurements fromthe light detector for the three light sources. By combining signals, itis possible to separate out the effects of changes in the oxygen contentof the tissue or changes in blood volume due to a change in bloodpressure from motion artifacts or muscle motion.

In another embodiment, the third light source is used to remove largechanges in the signals of the first and second light sources that aredue to mechanical artifacts. In some cases mechanical motion orvibration can cause rapid excursions of the optical signals that aremany times larger than those caused by a physiologic change inperfusion. These large excursions can be eliminated using the signalsfrom the third light source. This reduces the effects of the noise andfacilitates accurate determination of the change in perfusion. Ingeneral, the excursions created by the noise are faster than therelatively slow changes caused by physiologic perfusion changes.Therefore, the information about the excursions that is contained in thesignal from the third light source may be used to eliminate the noisefrom the first and second light sources using a suitable noiseelimination technique. The corrected first and second light source canthen be used as described above to determine tissue perfusion.

The tissue perfusion sensing system and method is particularlyapplicable to implantable medical devices. In particular, because thesystem and method can be implemented in a relatively simple device withrelatively simple processing it lends itself to applications such asimplantable cardiac devices. As one specific example, it is particularapplicable to implantable cardiac defibrillators that are designed to beinserted into the subcutaneous tissue of the patient and do not includeintravenous leads into the heart for sensing and defibrillation.

FIG. 1 is a schematic diagram of electronic circuitry included in amedical device according to an embodiment of the invention. Asillustrated in FIG. 1, a device 14 according to an embodiment of theinvention includes both a low voltage battery 153 and a high voltagebattery 112, for example, positioned within a hermetically sealedhousing (not shown) of the device 14. Low voltage battery 153 is coupledto a power supply (not shown) that supplies power to the devicecircuitry and the pacing output capacitors to supply pacing energy in amanner well known in the art. The low voltage battery 153 can includeone or more conventional LiCF_(x), LiMnO₂ or Lil₂ cells, while the highvoltage battery 112 can include one or more conventional LiSVO or LiMnO₂cells. It is understood that although the exemplary embodiment of FIG. 1includes both low and high power therapy, the present invention may beemployed in a device that provides only one therapy, such as a highpower defibrillation therapy, for example.

Device 14 functions are controlled by means of software, firmware andhardware that cooperatively monitor the ECG, determine when acardioversion-defibrillation shock or pacing is necessary, and deliverprescribed cardioversion-defibrillation and pacing therapies. FIG. 1incorporates circuitry set forth in commonly assigned U.S. Pat. Nos.5,163,427 “Apparatus for Delivering Single and Multiple Cardioversionand Defibrillation Pulses” to Keimel and 5,188,105 “Apparatus and Methodfor Treating a Tachyarrhythmia” to Keimel for selectively deliveringsingle phase, simultaneous biphasic and sequential biphasiccardioversion-defibrillation shocks, incorporated herein by reference intheir entireties

In FIG. 1, a sense amp 190 in conjunction with pacer/device timingcircuit 178 processes the far field ECG sense signal that is developedacross a particular ECG sense vector defined by a selected pair ofsubcutaneous electrodes 28 or, optionally, a virtual signal if selected.The selection of the sensing electrode pair is made through the switchmatrix/MUX 191 in a manner to provide the most reliable sensing of theEGM signal of interest, which would be the R wave for patients who arebelieved to be at risk of ventricular fibrillation leading to suddendeath. The far field ECG signals are passed through the switchmatrix/MUX 191 to the input of the sense amplifier 190 that, inconjunction with pacer/device timing circuit 178, evaluates the sensedEGM. Bradycardia, or asystole, is typically determined by an escapeinterval timer within the pacer timing circuit 178 and/or the controlcircuit 144. Pace Trigger signals are applied to the pacing pulsegenerator 192 generating pacing stimulation when the interval betweensuccessive R-waves exceeds the escape interval. Bradycardia pacing isoften temporarily provided to maintain cardiac output after delivery ofa cardioversion-defibrillation shock that may cause the heart to slowlybeat as it recovers back to normal function. Sensing subcutaneous farfield signals in the presence of noise may be aided by the use ofappropriate denial and extensible accommodation periods as described inU.S. Pat. No. 6,236,882 “Noise Rejection for Monitoring ECGs” to Lee, etal and incorporated herein by reference in its' entirety.

Detection of a malignant tachyarrhythmia is determined in the controlcircuit 144, for example, as a function of the intervals between R-wavesense event signals that are output from the pacer/device timing 178 andsense amplifier circuit 190 to the timing and control circuit 144.

Supplemental sensors such as tissue color, tissue oxygenation,respiration, patient activity are used to contribute to the decision toapply or withhold a defibrillation therapy as described in detail below.In particular, the invention includes optical sensors 117 to provide asecondary confirmation of a detected tachyarrhythmia event detected bythe device 14 by determining whether the heart is hemodynamicallyunstable in response to a tachycardia event being identified by thedevice 15 in response to R-wave sense intervals determined in theprimary detection algorithm, described below in detail. Sensorprocessing unit 194 provides sensor data to microprocessor 142 via databus 146. In addition to optical sensor 117, an activity sensor 119 mayalso be utilized so that patient activity and/or posture may also bedetermined by the apparatus and method as described in U.S. Pat. No.5,593,431 “Medical Service Employing Multiple DC Accelerometers forPatient Activity and Posture Sensing and Method” to Sheldon andincorporated herein by reference in its entirety. Similarly, patientrespiration may be determined by the apparatus and method as describedin U.S. Pat. No. 4,567,892 “Implantable Cardiac Pacemaker” to Plicchi,et al and incorporated herein by reference in its entirety. Opticalsensors 117 may be located on the housing of device 14, or may belocated on a lead.

Certain steps in the performance of the detection algorithm criteria arecooperatively performed in microcomputer 142, including microprocessor,RAM and ROM, associated circuitry, and stored detection criteria thatmay be programmed into RAM via a telemetry interface (not shown)conventional in the art. Data and commands are exchanged betweenmicrocomputer 142 and timing and control circuit 144, pacertiming/amplifier circuit 178, and high voltage output circuit 140 via abi-directional data/control bus 146. The pacer timing/amplifier circuit178 and the control circuit 144 are clocked at a slow clock rate. Themicrocomputer 142 is normally asleep, but is awakened and operated by afast clock by interrupts developed by each R-wave sense event, onreceipt of a downlink telemetry programming instruction or upon deliveryof cardiac pacing pulses to perform any necessary mathematicalcalculations, to perform tachycardia and fibrillation detectionprocedures, and to update the time intervals monitored and controlled bythe timers in pacer/device timing circuitry 178.

The algorithms and functions of the microcomputer 142 and controlcircuit 144 employed and performed in detection of tachyarrhythmias areset forth, for example, in commonly assigned U.S. Pat. Nos. 5,354,316“Method and Apparatus for Detection and Treatment of Tachycardia andFibrillation” to Keimel; 5,545,186 “Prioritized Rule Based Method andApparatus for Diagnosis and Treatment of Arrhythmias” to Olson, et al,5,855,593 “Prioritized Rule Based Method and Apparatus for Diagnosis andTreatment of Arrhythmias” to Olson, et al and 5,193,535 “Method andApparatus for Discrimination of Ventricular Tachycardia from VentricularFibrillation and Treatment Thereof” to Bardy, et al, (all incorporatedherein by reference in their entireties). Particular algorithms fordetection of ventricular fibrillation and malignant ventriculartachycardias can be selected from among the comprehensive algorithms fordistinguishing atrial and ventricular tachyarrhythmias from one anotherand from high rate sinus rhythms that are set forth in the '316, '186,'593 and '593 patents.

The detection algorithms are highly sensitive and specific for thepresence or absence of life threatening ventricular arrhythmias, e.g.,ventricular tachycardia (VT) and ventricular fibrillation (VF). When amalignant tachycardia is detected, high voltage capacitors 156, 158,160, and 162 are charged to a pre-programmed voltage level by ahigh-voltage charging circuit 164. It is generally consideredinefficient to maintain a constant charge on the high voltage outputcapacitors 156, 158, 160, 162. Instead, charging is initiated whencontrol circuit 144 issues a high voltage charge command HVCHG deliveredon line 145 to high voltage charge circuit 164 and charging iscontrolled by means of bi-directional control/data bus 166 and afeedback signal VCAP from the HV output circuit 140. High voltage outputcapacitors 156, 158, 160 and 162 may be of film, aluminum electrolyticor wet tantalum construction.

The negative terminal of high voltage battery 112 is directly coupled tosystem ground. Switch circuit 114 is normally open so that the positiveterminal of high voltage battery 112 is disconnected from the positivepower input of the high voltage charge circuit 164. The high voltagecharge command HVCHG is also conducted via conductor 149 to the controlinput of switch circuit 114, and switch circuit 114 closes in responseto connect positive high voltage battery voltage EXT B+ to the positivepower input of high voltage charge circuit 164. Switch circuit 114 maybe, for example, a field effect transistor (FET) with itssource-to-drain path interrupting the EXT B+ conductor 118 and its gatereceiving the HVCHG signal on conductor 145. High voltage charge circuit164 is thereby rendered ready to begin charging the high voltage outputcapacitors 156, 158, 160, and 162 with charging current from highvoltage battery 112.

High voltage output capacitors 156, 158, 160, and 162 may be charged tovery high voltages, e.g., 700-3150V, to be discharged through the bodyand heart between the electrode pair of subcutaneouscardioversion-defibrillation electrodes 113 and 123. The details of thevoltage charging circuitry are also not deemed to be critical withregard to practicing the present invention; one high voltage chargingcircuit believed to be suitable for the purposes of the presentinvention is disclosed. High voltage capacitors 156, 158, 160 and 162are charged by high voltage charge circuit 164 and a high frequency,high-voltage transformer 168 as described in detail in commonly assignedU.S. Pat. No. 4,548,209 “Energy Converter for Implantable Cardioverter”to Wielders, et al. Proper charging polarities are maintained by diodes170, 172, 174 and 176 interconnecting the output windings ofhigh-voltage transformer 168 and the capacitors 156, 158, 160, and 162.As noted above, the state of capacitor charge is monitored by circuitrywithin the high voltage output circuit 140 that provides a VCAP,feedback signal indicative of the voltage to the timing and controlcircuit 144. Timing and control circuit 144 terminates the high voltagecharge command HVCHG when the VCAP signal matches the programmedcapacitor output voltage, i.e., the cardioversion-defibrillation peakshock voltage.

Control circuit 144 then develops first and second control signalsNPULSE 1 and NPULSE 2, respectively, that are applied to the highvoltage output circuit 140 for triggering the delivery of cardiovertingor defibrillating shocks. In particular, the NPULSE 1 signal triggersdischarge of the first capacitor bank, comprising capacitors 156 and158. The NPULSE 2 signal triggers discharge of the first capacitor bankand a second capacitor bank, comprising capacitors 160 and 162. It ispossible to select between a plurality of output pulse regimes simply bymodifying the number and time order of assertion of the NPULSE 1 andNPULSE 2 signals. The NPULSE 1 signals and NPULSE 2 signals may beprovided sequentially, simultaneously or individually. In this way,control circuitry 144 serves to control operation of the high voltageoutput stage 140, which delivers high energycardioversion-defibrillation shocks between the pair of thecardioversion-defibrillation electrodes 113 and 123 coupled to the HV-1and COMMON output as shown in FIG. 1.

Thus, device 14 monitors the patient's cardiac status and initiates thedelivery of a cardioversion-defibrillation shock through thecardioversion-defibrillation electrodes 113 and 123 in response todetection of a tachyarrhythmia requiring cardioversion-defibrillation.The high HVCHG signal causes the high voltage battery 112 to beconnected through the switch circuit 114 with the high voltage chargecircuit 164 and the charging of output capacitors 156, 158, 160, and 162to commence. Charging continues until the programmed charge voltage isreflected by the VCAP signal, at which point control and timing circuit144 sets the HVCHG signal low terminating charging and opening switchcircuit 114. Typically, the charging cycle takes only fifteen to twentyseconds, and occurs very infrequently. The device 14 can be programmedto attempt to deliver cardioversion shocks to the heart in the mannersdescribed above in timed synchrony with a detected R-wave or can beprogrammed or fabricated to deliver defibrillation shocks to the heartin the manners described above without attempting to synchronize thedelivery to a detected R-wave. Episode data related to the detection ofthe tachyarrhythmia and delivery of the cardioversion-defibrillationshock may be stored in RAM for uplink telemetry transmission to anexternal programmer as is well known in the art to facilitate indiagnosis of the patient's cardiac state. A patient receiving the device14 on a prophylactic basis would be instructed to report each suchepisode to the attending physician for further evaluation of thepatient's condition and assessment for the need for implantation of amore sophisticated implantable cardio-defibrillator device (ICD). Inother embodiments, no storage of episode data will take place.

Turning now to FIG. 1A, a tissue perfusion sensor system 100 isillustrated schematically. The sensor system 100 includes at least twolight sources 102, a light detector 104, and a sensor controller 106.The first light source provides light at a wavelength where lightabsorption in the tissue is dependent upon the oxygen content of thetissue as well as the total content of venous and arterial blood in thetissue. The second light source provides light at a wavelength wherelight absorption in the tissue is substantially independent of theoxygen content in the blood, but where the light absorption is dependentupon the blood volume in the tissue. This is generally referred to as an“isobestic” wavelength. Light from the first and second light sourcesare emitted into the surrounding tissue and received back at the lightdetector after transmitting through, and/or being reflected by thesurrounding tissue. The sensor controller 106 receives the lightmeasurements from the light detector and uses the measurements at theisobestic light source to calculate the volume index, which indicatesthe change from a baseline value in the percentage of the volume of theilluminated tissue that contains blood. The controller also uses themeasurement from the first light source and the volume index tocalculate the oxygen index, which indicates the change from a baselinein the percentage of hemoglobin in the illuminated volume that is fullyoxygenated (oxyhemoglobin).

Specifically, a change in the measurement of first received lightcorresponds to a change in the overall oxygen content of the tissue aswell as a change in blood volume of the tissue. Therefore, that changecan be the result of a drastic change in perfusion, such as during afibrillation event when perfusion stops altogether and local tissueoxygenation starts falling due to on-going metabolic activity and lackof oxygen replacement, as well as changes in the volume of arterial andvenous blood due to a drop in arterial pressure or due tovasoconstriction or vasodilatation. The first received light is alsodependent on changes in blood volume due to posture, compression oftissue or tissue motion. Thus, by itself, the received first lightmeasurement cannot be used to reliably detect a change in the tissueperfusion. However, because the second light source was chosen to have awavelength where absorption in the tissue is independent of oxygencontent in the blood, but is dependent upon blood volume, the sensorcontroller 106 can use the second light measurement to determine theamount of the blood volume changed from a baseline. This change may havebeen due to the effects of posture, muscle motion or a drop in bloodpressure. The effects of the change in blood volume can then be removedfrom the measurements of the first light source in order to determinethe amount of change in the tissue oxygenation. As one example, thesensor controller 106 scales the change from a baseline of the secondlight measurement by a gain constant and then subtracts the scaledchange in the second measurement from the change from a baseline of thefirst light measurement. This results in a compensated value that willtrack a change in tissue oxygenation substantially independent of bloodvolume changes. Thus, the sensor controller 106 determines if a changein the optical properties of tissue are due to a change in the bloodvolume, the tissue oxygenation, or both. The volume index and theoxygenation index may be used individually or in combination to makedecisions on the status of the patient. For example, a change in thevolume index, indicative if a drop in the blood volume, that isassociated with a similar change in the tissue perfusion index,indicating a drop in tissue oxygenation, may be interpreted as a loss ofperfusion due to cardiac fibrillation and the decision may be made todefibrillate. If there is no drop in oxygenation, the decision may bethat the drop in volume was associated with a change in posture orcompression of the tissue and defibrillation is not needed.

Turning now to FIG. 2, a first embodiment of tissue perfusion sensorsystem 200 is illustrated schematically. The tissue perfusion sensorsystem 200 includes a first light source 202, a second light source 204,and a light detector 206. The tissue perfusion sensor system 200 isformed in a housing 216, which includes a biocompatible opticallytransparent window 212 and a biocompatible optically transparent polymercover 214. The first light source 202, second light source 204 and lightdetector 206 are attached to a substrate 208, such as a ceramicsubstrate. The detector 206 is blocked off from the first light source202 and second light source 204 by an optical barrier 210.

The tissue perfusion sensor 200 would typically be implemented as partof an implantable medical device, and the housing 216 could suitably bepart of the housing for the larger device or separate and distinct. Theimplantable medical device can be a variety of different types ofdevices, such as an implantable cardiac device. Even more specifically,the tissue perfusion sensor 200 can be implemented as part of animplantable cardiac defibrillator that is designed to be inserted intothe subcutaneous tissue of the patient. In such an application theimplantable device would typically be surgically implanted in thesubcutaneous chest area.

During operation, the first light source 202 outputs light at a firstwavelength, and the second light source 204 outputs light at a secondwavelength. The light from the first and second light sources passesthrough the optical window 212 and polymer cover 214 and into the nearbytissue 220. A portion of the light from those sources is reflected backfrom the nearby issue 220, through and into the detector 206.

A variety of different types of devices can be used for the first andsecond light sources. As one example, the first and second light sourcescan each comprise light emitting diodes (LEDs). LEDs are commonly usedlight sources that provide light in a relatively narrow wavelength band.It should also be noted that a single wide-band light source such as awhite LED can be used to provide all the wavelengths needed to performthe sensing operation. It should be noted however, that a singlewide-band source would require either multiple detectors each with anarrow-optical filter or a grating that would separate the light outinto a spectrum that falls on a detector array.

Lasers are another type of device that can be used as the first andsecond light sources. Lasers provide several advantages over LEDs,including having an even narrower optical spectrum, a narrower opticaldivergence angle, and a more efficient conversion of electrical powerinto optical power. Efficiency and narrower beam divergence areimportant in an implantable device field since both will help reduceenergy consumption of the sensor. The narrow band aspect of thesemiconductor laser is also desirable in that it enables measurement ata very narrow wavelength range where tissue properties can be studied indetail and it also gives the option of filtering the returning light tosuppress the undesired effects of broadband ambient light.

One particular type of laser that is suitable is a vertical cavitysurface emitting laser (VCSEL). In addition to the advantages providedby all lasers, the VCSELs also have the advantage of emitting the lightout of the top surface. This permits simple design of the opticalpackage and easy transmission of the light out of the sensor.

Finally, it should be pointed out that different types of light sourcescan be used in one implantable medical device. For example, acombination of LEDs and VCSELs can be used to provide the desired lightsources.

As described above, the first light source 202 provides light at awavelength where light absorption in the tissue is dependent upon theoxygen content of blood in the tissue and the blood volume. The secondlight source 204 provides light at a wavelength where light absorptionin the tissue is substantially independent of the oxygen content in theblood, but where the light absorption is dependent upon the blood volumein the tissue. This is generally referred to as an “isobestic”wavelength.

As one example, the first light source provides light at a wavelength ofapproximately 660 nm, in the red region. The second light sourceprovides light at a wavelength at approximately 805 nm, in the infrared(IR) region. It should be noted that in this example the light sourceswould typically output light in a wavelength band surrounding 660 nm and805 nm respectively, and that the width of those wavelength bands wouldagain depend on the type of light sources being used. Thus, in oneembodiment the first light source could provide light having awavelength anywhere between 650 nm and 670 nm, and the second lightsource could provide light having a wavelength anywhere between 695 nmand 830 nm.

Light from the first and second light sources is emitted into thesurrounding tissue and received back at the light detector 206 afterbeing reflected by the nearby tissue 220. A variety of different typesof devices can be used for the light detector 206. For example, a photodetector or array of photo detectors can be used. As will be describedin greater detail below, a single photo detector could be used if thelight at different wavelengths is emitted in a time-multiplexed mannerso that the reflected light can be measured via the same photo detectorat different instants of time. In another embodiment each light sourcewould fire at the same time and the light detector 206 would use morethan one photo detector, with each photo detector dedicated to aparticular wavelength of interest. In this embodiment each photodetector could include an optical filter that passes only the wavelengthit is trying to measure. In the case where a white LED is used as thelight sources, multiple photo detectors integrated with desired opticalfilters or a diffraction grating in combination with a linear photodetector array can be used to select optical response at the desiredwavelengths

In this embodiment, a sensor controller (not shown in FIG. 2) receivesthe light measurements from the light detector 206, the measurementscorresponding to light from the first light source 202 and second lightsource 204. In general, the controller calculates the tissue oxygenationindex and blood volume index in the surrounding tissue using thedetected first light and the detected second light. From this, thesensor controller determines if a change in the measurement of detectedfirst light source 202 is likely caused by a lack of perfusion, or if itinstead likely caused by other factors, such as muscle movement by thepatient. Specific examples of such a calculation will be discussed ingreater detail below.

It should also be noted that the sensor controller could be any type ofdevice, such as a suitable digital processor or application specificintegrated circuit (ASIC). It should also be noted that the sensorcontroller could be a separate processor, or could be part of processorused by the implantable medical device.

Turning now to FIG. 3, a second exemplary embodiment of a tissueperfusion sensor system 300 is illustrated schematically. The tissueperfusion sensor system 300 again includes a first light source 302, asecond light source 304, and a light detector 306. Additionally, thisembodiment includes a third light source 305. Furthermore, in thisembodiment the light sources are in housing 316, while the lightdetector 306 is separate in housing 315. This allows the light to betransmitted through the tissue 320 located between the light sources302, 304, and 305 and the light detector 306, instead of using reflectedlight as in system 200.

Like system 200, the tissue sensor 300 would typically be implemented aspart of an implantable medical device. In this case either housing 315or 316 could be part of the housing for the larger device, or they couldeach be separate and distinct.

In this embodiment, the first light source 302 outputs light at a firstwavelength, the second light source 304 outputs light at a secondwavelength, and third light source 305 outputs at a third wavelength.The light from these sources passes through the tissue 320 and into thedetector 306. Again, a variety of different types of devices can be usedfor the first, second and third light sources, including LEDs andlasers.

As described above, the received light from all three light sources 302,304 and 305 are dependent on the volume of blood in the tissue 320either due to changes in arterial pressure, motion orvasoconstriction/dilation. The received light signals from the firstlight source 302 and the third light source 305 are also dependent uponthe oxygen content of the tissue 320. The second light source 304provides light at a wavelength where light absorption in the tissue 320is substantially independent of the oxygen content in the tissue 320.

As one example, the first light source 302 provides light having awavelength centered around approximately 660 nm, the second light source304 provides light at a wavelength centered around approximately 800 nm,and the third light source 305 provides light at a wavelength centeredaround approximately 910 nm. It should be noted that the first and thirdlight sources can interchangeable, and thus the first light source couldcomprise 910 nm and the second light source could comprise 660 nm. Thesensor controller (not shown in FIG. 3) receives the light measurementsfrom the light detector 306 for light from each of the three sources.Again, the sensor controller combines the measurements from the threelight sources to determine the degree to which tissue oxygen ischanging, and further determine if the blood volume is changing due to aloss of arterial pressure or other factors, such as muscle movement bythe patient. Thus, when combined with measurements from the first lightsource 302 and the second light source 304, the addition of a thirdlight source 305 gives the system and method the ability to increase thereliability of the perfusion detection. Additionally, by choosingwavelengths for the first and third light sources such that one isshorter and one is longer than the second (isobestic) wavelength,perfusion detection specificity can be enhanced by relying on the factthat detected signals from the first and third light sources will changein opposite directions due to changes in tissue oxygen.

Additionally, the measurements of the third light source can be used tocompensate for the effects of mechanical vibration and other noise inthe measurements of the first and second light sources. Specifically,the measurements of the third light source can be used to remove largeexcursions in the measurements that can be caused by mechanicalvibration and noise. In some cases these large excursions can be manytimes larger than the changes in the measurements caused by physiologicchanges in perfusion. The measurements from the third light source canbe used to reduce the effects of the excursions on the measurements fromthe first and second light sources, and can thus be used to facilitate amore accurate determination of the change in perfusion from the firstand second light sources.

Turning now to FIG. 4, a flow diagram 400 illustrates an exemplarymethod for detecting a change in perfusion by calculating changes intissue oxygenation and blood volume. In this example, the decision aboutthe change in perfusion is based on the changes that occur in the bloodvolume and tissue oxygen over a few seconds time span. In order to reacha quality decision, the optical signals are corrected for thedetrimental effects of ambient light, tissue encapsulation, and motionartifacts. These detected changes in perfusion can then be used by animplantable medical device to determine what action, if any, to take.

The method 400 begins when the medical device requests input regardingtissue perfusion. This could occur for a variety of reasons. As oneexample, an implantable cardiac device can use a calculation ofperfusion to determine whether or not defibrillation is desirable.Specifically, when an ECG based algorithm suspects a VF or untoleratedVT episode, the method 400 can be used to determine if instead thepatient is experiencing a tolerated VT episode where defibrillationwould not be desirable. During fibrillation, both the tissue oxygen andthe blood volume change over the course of many seconds. The measure oftissue oxygen and blood volume as a function of time can be used todetermine whether the patient is in fibrillation Of course, this is justone example of how a medical device, including external and implantablemedical devices, can use a determination of change in tissue perfusionto determine delivery of therapy.

According to an embodiment of the invention, raw optical signals arecorrected for errors due to ambient light, tissue encapsulation, andmechanical vibration. Light levels are measured at background levels andat least two specific wavelengths, Block 402. The measurements of thebackground light level with the optical sources turned off are used todetermine the current levels of ambient light, which is then utilized tocompensate the measurements at the two or more wavelengths for anypotential offset in the measured data caused by the ambient light, Block404. The measurements at two or more wavelengths will also be utilizedto calculate changes in the tissue oxygenation index and changes intissue blood volume index, as described below.

As one specific example, if the three light sources 302, 304 and 305 areutilized, five distinct measurements of light are taken, including, afirst ambient light measurement D₁ taken at ambient light levels, alight measurement L₁ taken in response to light from the first lightsource 302, a light measurement L₂ taken in response to light from thesecond light source 304, a light measurement L₃ taken in response tolight from the third light source 305, and a second ambient lightmeasurement D₂ taken at ambient light levels. As described below, byutilizing ambient light measurements, the device is able to moreaccurately correct the light measurements from the three light sources302, 304 and 305 for the effects of ambient light.

The first light source 302 emits light at a first wavelength into thesurrounding tissue, and the light is received back at the detector 306after having been reflected by, and/or passing through, the surroundingtissue so that the first light measurement L₁ is a measurement of thelevel of light received by the detector 306 from the first light source302. Alternatively, light measurement L₁ can correspond to the timeinterval needed to integrate the amount of received light to reach adefined threshold. In both cases, light measurement L₁ is a measurementof light that is responsive to reflectivity of the surrounding tissue.

Likewise, the second light source 304 is used to provide light at asecond wavelength, and the third light source 305 is used to providelight at a third wavelength. Similar to the first light source 302, thelight from these light sources 304 and 305 is also emitted into thesurrounding tissue and received back at the detector 306, so that lightmeasurements L₂ and L₃ are measurements of the second and third lightsources 304 and 305.

The three light measurements L₁, L₂ and L₃ are taken at threewavelengths that are selected based upon tissue absorption.Specifically, the first wavelength is selected such that absorption inthe tissue is dependent upon the oxygen content and the volume of bloodin the tissue. The second light wavelength is selected to be isobestic,where absorption is substantially independent of the oxygen content inthe tissue, but where the light absorption is dependent upon the bloodvolume in the tissue. The third wavelength is also selected such thatlight absorption in the tissue is dependent upon the oxygen content andblood volume in the tissue. Additionally, the first light wavelength isselected to be shorter than the second light wavelength, and the thirdlight wavelength is selected to be longer than the second lightwavelength.

FIG. 4A is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention. As illustrated in FIGS. 3 and 4A, during the adjusting of thelight intensities measured via the light sources for the effects ofambient light (Block 404 of FIG. 4), the device measures the backgroundlight level by measuring the light detected at detector 306 with thethree light sources 302, 304 and 305 turned off to generate a firstambient light measurement D₁, Block 430. The device then determines theresulting light detected at detector 306 at each of the predeterminedlight sources, 432. For example, the device determines the resultinglight detected at detector 306 from only the first light source 302 togenerate a first light measurement made at the first wavelength L₁, andthen determines whether a light measurement has been made at allwavelengths, Block 434. A determination of the resulting light detectedat detector 306 from only the second light source 304 is then maderesulting in a second light measurement being made at the secondwavelength L₂, Block 432, and a determination as to the resulting lightdetected at detector 306 from only the third light source 305 is made togenerate a third light measurement made at the third wavelength L₃,Block 432.

Once the three light measurements L₁, L₂, and L₃ have been generated,Yes in Block 434, the device disables the three light sources 302, 304and 305, and measures the ambient light detected at detector 306 togenerate a second ambient light measurement D₂, Block 436.

Using the generated light measurements at different wavelengths L₁, L₂,and L₃ and at ambient light D₁, and D₂, the device then corrects each ofthe light measurements at the three different wavelengths for theeffects of ambient light, Block 438. For example, according to anembodiment of the invention, a corrected light measurement made at thefirst wavelength L₁′, along with a corrected light measurement made atthe second wavelength L₂′ and a corrected light measurement made at thethird wavelength L₃′ is generated using Equation 1:

$\begin{matrix}{{L_{1}^{\prime} = {L_{1} - \left( {{\frac{1}{4}D_{2}} + {\frac{3}{4}D_{1}}} \right)}}{L_{2}^{\prime} = {L_{2} - \left( {{\frac{1}{2}D_{1}} + {\frac{1}{2}D_{2}}} \right)}}{L_{3}^{\prime} = {L_{3} - \left( {{\frac{3}{4}D_{2}} + {\frac{1}{4}D_{1}}} \right)}}} & {{Equation}\mspace{20mu} 1}\end{matrix}$

As can be seen, using Equation 1, the light measurements are scaled bythe ambient light measurements based in part on the relative timing ofthe measurements. Thus, the first light measurement made at the firstwavelength L₁ is scaled more heavily using the first ambient lightmeasurement D₁ than the second ambient light measurement D₂, given therelative time proximity between the measurements. Conversely, the thirdlight measurement made at the third wavelength L₃ is scaled more heavilyusing the second ambient light measurement D₂ than the first ambientlight measurement D₁. Finally, the second light measurement made at thesecond wavelength L₂ is scaled evenly by first ambient light measurementD₁ and the second ambient light measurement D₂ based on the fact thatthe second light measurement made at the second wavelength L₂ issubstantially centered between the two ambient light measurements D₁ andD₂.

According to another embodiment of the invention, once all of three ofthe light measurements at three different wavelengths and the twoambient light measurements have been generated, rather than utilizingEquation 1 to generate the corrected light measurements L₁′, L₂′ andL₃′, the device corrects each of the three light measurements at thethree different wavelengths for the effects of ambient light, Block 438,by subtracting the average of the two ambient light measurements D₁ andD₂ from each of the light measurements L₁, L₂ and L₃.

FIG. 4B is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention. As illustrated in FIG. 4B, according to an embodiment of theinvention, in order to reduce the effects of ambient light on the lightsource during generation of a light measurement, Block 404 of FIG. 4, anambient light measurement is made prior to and after each lightwavelength measurement L₁, L₂ and L₃, which is then used to generate acorresponding corrected light measurement. For example, the devicemeasures the light detected at detector 306 with the three light sources302, 304 and 305 turned off, Block 440, to generate a first ambientlight measurement D₁. The device then measures the resulting lightdetected at detector 306 from only the first light source 302 togenerate a first light measurement made at the first wavelength L₁,Block 442, and subsequently again measures the light detected atdetector 306 with the three light sources 302, 304 and 305 turned off,Block 444, to generate a second ambient light measurement D₂.

The process continues with the controller generating the second lightmeasurement made at the second wavelength L₂, subsequent to thegeneration of the second ambient light measurement D₂, and againmeasuring the light detected at detector 306 with the three lightsources 302, 304 and 305 turned off, Block 444, to generate a thirdambient light measurement D₃. A third light measurement of lightgenerated by the third light source at the third wavelength L₃ isperformed subsequent to the generation of the third ambient lightmeasurement D₃, and once again the light detected at detector 306 withthe three light sources 302, 304 and 305 turned off is measured, Block444, to generate a fourth ambient light measurement D₄.

Once all of the light measurements associated with the predeterminednumber of light sources have been generated, Yes in Block 446, thedevice then corrects each of the light measurements at differentwavelengths for the effects of ambient light, Block 448, using theambient light measurement generated prior to and subsequent to thegenerated light measurement. For example, according to an embodiment ofthe invention, a corrected light measurement made at the firstwavelength L₁′ is generated by subtracting the average of the firstambient light measurement D₁ and the second ambient light measurement D₂from the first light measurement made at the first wavelength L₁.Similarly, a corrected light measurement made at the second wavelengthL₂′ is generated by subtracting the average of the second ambient lightmeasurement D₂ and the third ambient light measurement D₃ from thesecond light measurement made at the second wavelength L₂, and acorrected light measurement made at the third wavelength L₃′ isgenerated by subtracting the average of the third ambient lightmeasurement D₃ and the fourth ambient light measurement D₄ from thethird light measurement made at the third wavelength L₃.

FIG. 4C is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention. As illustrated in FIG. 4C, according to an embodiment of theinvention, in order to reduce the effects of ambient light on the lightsource during generation of a light measurement, Block 404 of FIG. 4,the device determines the light detected at detector 306 with the threelight sources 302, 304 and 305 turned off, Block 450, to generate afirst ambient light measurement D₁. The device then determines theresulting light detected at detector 306 from only the first lightsource 302 to generate a first light measurement made at the firstwavelength L₁, Block 452, and subsequently again determines the lightdetected at detector 306 with the three light sources 302, 304 and 305turned off, Block 454, to generate a second ambient light measurementD₂.

A determination is then made as to whether one of the measured firstambient light measurement D₁ and the second ambient light measurementD₂, or an average of the measured first ambient light measurement D₁ andthe second ambient light measurement D₂ is greater than an ambientthreshold, Block 456. The ambient threshold may correspond to apredetermined fixed value, such as one half of a dynamic rangeassociated with the measurement of the light intensity, for example. Ifthe ambient light is determined to be greater than the ambientthreshold, Yes in Block 456, the device adjusts the light emitted fromthe first light source 302, Block 458. For example, the device adjuststhe output of the light source 302 so that the amplitude of the lightsignal emitted is increased and the pulse width is decreased. In oneembodiment, the light signal is adjusted so that the amplitude isdoubled and the length of time that the light is measured is reduced byone half. Once the light emitted by the first light source 302 isadjusted, Block 458, the device repeats the determination of the lightdetected from the first light source 302, Blocks 450-454 using theadjusted light emitting settings.

If the ambient light is determined to be less than or equal to theambient threshold, No in Block 456, the current generated lightmeasurement and the ambient light measurements immediately preceding andsubsequent to the current generated light measurement are stored, andthe process is repeated for the remaining light sources. For example,once one of the measured first ambient light measurement D₁ and thesecond ambient light measurement D₂, or an average of the measured firstambient light measurement D₁ and the second ambient light measurement D₂is less than or equal to the ambient threshold, No in Block 456, thedevice then determines the resulting light detected at detector 306 fromonly the second light source 304 to generate a second light measurementmade at the second wavelength L₂, Block 452, and subsequently againdetermines the light detected at detector 306 with the three lightsources 302, 304 and 305 turned off, Block 454, to generate a thirdambient light measurement D₃.

A determination is then made as to whether the determined third ambientlight measurement D₃ is greater than the ambient threshold, Block 456.If the third ambient light measurement D₃ is greater than the ambientthreshold, the device adjusts the light emitted from the second lightsource 304, Block 458, as described above, and repeats the determinationof the light detected from the second light source 304, Blocks 450-454using the adjusted light emitting settings. Once the third ambient lightmeasurement D₃ is determined to be less than or equal to the ambientthreshold, No in Block 456, the second light measurement made at thesecond wavelength L₂ and the third ambient light measurement D₃ arestored, and the process is repeated for the third light source 305. Inparticular, the device then determines the resulting light detected atdetector 306 from only the third light source 305 to generate a thirdlight measurement made at the third wavelength L₃, Block 452, andsubsequently again determines the light detected at detector 306 withthe three light sources 302, 304 and 305 turned off, Block 454, togenerate a fourth ambient light measurement D₄.

A determination is then made as to whether the measured fourth ambientlight measurement D₄ is greater than the ambient threshold, Block 456.If the fourth ambient light measurement D₄ is greater than the ambientthreshold, the device adjusts the light emitted from the third lightsource 304, Block 458, as described above, and repeats the determinationof the light detected from the third light source 305, Blocks 450-454using the adjusted light emitting settings. Once the fourth ambientlight measurement D₄ is determined to be less than or equal to theambient threshold, No in Block 456, the third light measurement made atthe second wavelength L₃ and the fourth ambient light measurement D₄ arestored. Once the three light measurements have been generated, Yes inBlock 460, the device then corrects each of the three light measurementsat different wavelengths for the effects of ambient light, Block 462,using the ambient light measurement generated prior to and subsequent toeach of the generated light measurements, and one or a combination ofthe methods described above in reference to Block 438 or Block 448 ofrespective FIGS. 4A and 4.

FIG. 4D is a flowchart of a method of adjusting for ambient light duringsensing of signals in a medical device according to an embodiment of theinvention. In another embodiment, the device determines the lightdetected at detector 306 with the three light sources 302, 304 and 305turned off, Block 470, to generate a first ambient light measurement D₁.A determination is then made as to whether the first ambient lightmeasurement D₁ is less than a predetermined ambient light threshold,Block 472. The predetermined ambient light threshold may correspond to apercentage of an ambient light dynamic range, such as 10% of the ambientlight dynamic range for example. If the first ambient light measurementD₁ is greater than or equal to the predetermined ambient lightthreshold, No in Block 472, the controller device adjusts the detector306, Block 474, and the first ambient light measurement D₁ is generatedagain, Block 470, using the adjusted settings. According to oneembodiment, the device adjusts the detector 306 by reducing the gain ofthe detector 306 by a predetermined amount, such as by 2 for example.The process is repeated so that once the first ambient light measurementD₁ is less than the predetermined ambient light threshold, Yes in Block472, the controller device determines the resulting light detected atdetector 306 from only the first light source 302 to generate a firstlight measurement made at the first wavelength L₁, Block 476. Adetermination is then made as to whether the output corresponding to thefirst light measurement made at the first wavelength L₁ is greater thana predetermined upper limit threshold, Block 478. If the first lightmeasurement made at the first wavelength L₁ is greater than thepredetermined upper limit threshold, the controller device adjusts thefirst light source 302, Block 480, and the first light measurement madeat the first wavelength L₁ is generated again, Block 476, and theprocess is repeated. According to one embodiment, in order to adjust thefirst light source 302 when the first light measurement made at thefirst wavelength L₁ is greater than the predetermined upper limitthreshold, the device reduces the current of the light source 302 by apredetermined amount, for example. For LEDs, the current into the LEDmay be reduced by 50% of the current used for the previous measurement.For lasers the current into the laser may be reduced by 50% of thedifference between the previous current and the threshold current of thelaser.

Once the first light measurement made at the first wavelength L₁ is nolonger greater than the predetermined upper limit threshold, No in Block478, a determination is made as to whether the output corresponding tothe first light measurement made at the first wavelength L₁ is less thana predetermined lower limit threshold, Block 482. If the first lightmeasurement made at the first wavelength L₁ is less than thepredetermined lower limit threshold, the controller device adjusts thefirst light source 302, Block 484, the first light measurement made atthe first wavelength L₁ is generated again, Block 486, and thedetermination of whether the first light measurement made at the firstwavelength L₁ is less than the predetermined lower limit threshold isrepeated. According to one embodiment, in order to adjust the firstlight source 302 when the first light measurement made at the firstwavelength L₁ is less than the predetermined lower limit threshold, thedevice increases the current of the light source 302 by a predeterminedamount, for example. For LEDs, the current into the LED may be reducedby 50% of the current used for the previous measurement. For lasers thecurrent into the laser may be reduced by 50% of the difference betweenthe previous current and the threshold current of the laser.

Once either the initial or the adjusted first light measurement made atthe first wavelength L₁ is no longer less than the predetermined lowerlimit threshold, No in Block 482, the first ambient light measurement D₁and the first light measurement made at the first wavelength L₁ arestored, and the process is repeated for the remaining light detectors304 and 305. When the three light measurements have been generated, Yesin Block 488, the device then corrects each of the three lightmeasurements at different wavelengths for the effects of ambient light,Block 490, using the stored adjusted ambient light measurements and theadjusted light measurements and one or a combination of the methodsdescribed above in reference to Block 438 or Block 448 of respectiveFIGS. 4A and 4.

According to an embodiment of the invention, rather than adjusting eachof the outputs of the light sources using an upper and a lower limitthreshold, Blocks 478-486, the device may merely make a single currentadjustment to the light source to obtain a desired output, such as apercentage of predetermined dynamic range, such as 60% for example.

According to another embodiment, once the four ambient lightmeasurements have been obtained, the device uses the four ambient lightmeasurements to perform a cubic fit to generate an ambient light profileas a function of time that can then be used to predict the ambient lightvalues that are then utilized by the device to correct each of the threelight measurements at different wavelengths for the effects of ambientlight.

In addition to correcting the light measurements for the effects ofambient light, the device may also to correct the light measurements fortissue encapsulation around the light sources and light detectors, Block406. When a foreign device is implanted into the body a typical immuneresponse causes fibrous tissues to be formed around the device. Thefibrous tissues typically have a very low blood perfusion, and thepresence of the fibrous tissue can degrade the optical signal becauselight may enter the tissue encapsulation and reach the photodetectorwithout interacting with the perfused tissue. In addition, it can causeless light to reach the perfused tissue, resulting in less change (e.g.,gain) in the optical signal as the level of perfusion changes. Thus, thefibrous tissues can degrade the levels of received light and interferewith the ability to calculate the tissue oxygen content or blood volumeby causing a shift in the average value of light received and the amountof change as the perfusion changes. As described below, in order tocorrect the light measurements for tissue encapsulation, therebyimproving the accuracy of the corrected measurements, the deviceaccording to an embodiment of the invention utilizes a tissueencapsulation factor, Block 420, to correct the light measurements andcompensate for encapsulation.

Several different techniques could be used to generate the tissueencapsulation factor, Block 420. In one example, the tissueencapsulation factor would be determined at times during the day whenthe patient is stable and at rest. For example, the implantable medicaldevice can be configured to selectively emit and receive light at thevarious use intervals to calibrate and determine the tissueencapsulation factor. In this example, once the device determines thatit is time to update the encapsulation factor, signals from anaccelerometer and a heart rate monitor within the implanted device areevaluated to determine whether the patient is resting and has a normalheart rate. If the patient is not resting the system goes into a holdingpattern for a predetermined amount of time, e.g. one hour, and thentries again. If the patient is resting, then the optical sensor measuresand stores optical signals for a period long enough to contain a smallnumber of heart beats (e.g. 5 seconds) and at a sampling rate fastenough to resolve the swings in the optical signal caused by the pulsingof the arteries (e.g. 15 samples/sec.). The measurement of receivedlight during these periods of rest is then used to generate signalsrepresentative of the average value of light that is received and theamplitude of the swing in the light received due to the pulsing of thearteries. Those values can then be compared to a stored baseline valueor historical values to determine the effect that the tissueencapsulation has on the received optical signals. Typically, thesecalculations would be done more frequently right after implantation ofthe device because the healing process within the first few weeks afterimplant would cause changes in the optical signals at a relatively fastrate. Once the initial healing process is completed, the updates canoccur at less frequent intervals.

As an example, after the period of initial fast healing (e.g., 3 weeks),the device can be configured to calculate the factors once a day. Afterthis initial period, the device can be configured to calculate thefactor once every 10 days. An alternative is to allow the device todetermine the frequency of updates by tracking how much the correctionfactors change between updates. If the factors are changingsignificantly between updates, then the updates should be more frequent.As the variation in the correction factors between updates decreases,the updates can be made less frequent.

FIG. 4E is a flow chart of a method for correcting sensing by a medicaldevice for the effects of tissue encapsulation according to anembodiment of the invention. As illustrated in FIG. 4E, once the devicedetermines it is time to perform an updated calculation of normalizationcoefficients, Yes in Block 506, that will be utilized to generate thetissue encapsulation factor (Block 420 of FIG. 4), a determination ismade as to whether the patient is currently physically inactive, Block508. In the embodiment of FIG. 4E, it is desired that the tissueencapsulation factor be generated while the patient is at rest for atleast a predetermined time period, such as ten seconds for example, andthis determination can be made using a vibration sensor, such as anaccelerometer, for example. The patient may be determined to be inactiveif in either a prone position or an upright position for thepredetermined time period. According to one embodiment, thedetermination as to whether the patient is inactive may include adetermination of the time of day and the patient position so that thepatient is likely asleep, such as during the evening or early morninghours.

In the determination of whether an updated calculation of thenormalization coefficients is required, Block 506, the device initiallydetermines whether a predetermined healing period from the time that thesensors are initially implanted in the patient, such as 3 weeks forexample, has expired before the calculation of the normalizationcoefficients is initially generated. Once the healing period has ended,the device then begins periodically performing the updating of thenormalization coefficients once over predetermined time period apredetermined number of times, which are then subsequently updated. Forexample, according to one embodiment, once the device generates updatednormalization coefficients every two weeks for ten weeks, resulting inthe coefficients being updated five times. After ten weeks, the devicethen generates the coefficients once a month for nine months, resultingin the coefficients being updated another nine times, after which pointgeneration of new normalization coefficients is generated once every 2months for a one year time period, and so forth.

According to another embodiment, the amount of time between the updatingof the coefficients depends upon a comparison of the current generatednormalization coefficients to the previously generated coefficients. Forexample, once the normalization coefficients are initially generatedafter the three week healing period from the time of implant, resultingin a baseline being generated as described below, updated normalizationcoefficients are generated a predetermined time period later, such astwo weeks for example, resulting in a new baseline being generated thatis then compared to the initially generated baseline. Based upon thecomparison between the two baseline values, the time period after whichthe next updated coefficients are generated is determined. For example,if the difference between the current updated baseline is less than orequal to a predetermined threshold, indicating the effect of tissueencapsulation has not significantly changed since the last updating ofthe coefficients, the time period for generating the updatednormalization coefficients is updated, such as from every two week toevery four weeks, for example. However, if the difference between thecurrent updated baseline is greater than the predetermined threshold,indicating the effect of tissue encapsulation has significantly changedsince the last updating of the coefficients, the time period forgenerating the updated normalization coefficients is updated, such asfrom once every two week to once every one week, for example.

If the patient is determined not to be in the desired physical mode, thedevices waits a predetermined time period, Block 510, such as one hourfor example, and again makes the determination as to whether the patientis inactive, Block 508. Once the patient is determined to be in thedesired physical mode, Yes in Block 508, the device determines theresulting light detected at the detector to generate a light measurementfor each light source, Block 512, which are then stored, Block 514. Itis understood that the calculation of the normalization coefficients ismade separately for each light source, so that if three light sourcesare utilized, for example, the calculation is made for each of the threelight sources to generate respective normalization coefficients utilizedto correct the respective light measurements subsequently made for eachof the light sources for tissue encapsulation, Blocks 420 and 406 ofFIG. 4.

Since it is desirable that the light measurements be made at least overa completed cardiac cycle, the detecting of the resulting light detectedat the detector could begin at the detection of an R-wave and end on thenext detected R-wave. According to another embodiment, rather thanensuring the measurements are taken over a complete waveform, themeasurements could be correlated with peak-to-peak arterial pressurepulses. According to yet another embodiment, the measurements could betaken over a predetermined time period, such as one second for example.In each embodiment, the measurements are taken for each light source ata predetermined rate, such as 15 measurements over the cardiac orarterial pressure cycle, or at a rate of 15 measurements per second whenthe one second time period is utilized.

FIG. 4F is a graphical representation of the AC pulse amplitude and theDC average optical amplitude associated with the method for correctingsensing by a medical device for the effects of tissue encapsulation ofFIG. 4E according to one embodiment of the invention. As illustrated inFIG. 4F, the curve 527 represents measured amplitude of light as afunction of time, from one light source, that passes through perfusedtissue. The pulsing of the blood in the arteries and arterioles changesthe volume of blood in the tissue for every heart beat. The result isthat curve 527 has a pulsatile shape. Each pulse represents the timefrom one heart beat to the next, called the cardiac cycle 528. Eachcardiac cycle has a maximum value and a minimum value. The differencebetween the maximum and the minimum is the AC pulse amplitude, 523. Themeasured waveform 527 also has an average DC value, 525, which is theaverage amplitude of all of the points of curve 527 during the cardiaccycle 528.

Once light measurements are obtained for the desired cardiac or arterialpressure cycle, or the one second time period, Yes in Block 516, and theAC pulse amplitude and DC average optical amplitudes for the completecycle or time period have been determined for each light source, Block518, a determination is made as to whether light measurements have beenobtained for a predetermined number of cardiac cycles or time periods,Block 520. The number of cardiac cycles or time periods utilized inBlock 520 will depend upon what the user determines can be averaged toreduce the error in any one measurement and provide a valuerepresentative of the true AC and DC measures. According to oneembodiment, for example, the light measurements are taken over tencardiac cycles or ten one second time periods. Once the lightmeasurements have been obtained for the predetermined number of cardiacor arterial cycles, or for the one second time periods, No in Block 520,the device calculates, from the 10 stored AC and DC amplitudes, anaverage or median AC pulse amplitude and average or median DC amplitude,Block 518, for each of the wavelengths, Block 522.

Once the average AC pulse amplitude and the average DC amplitude aredetermined, Block 522, the device determines a gain coefficient for eachof the light sources, Block 524, using the average AC pulse amplitudedetermined in Block 522 and a determined baseline AC pulse amplitude,and an offset coefficient, Block 526 for each of the light sources,using the current average DC amplitude determined in Block 522 and abaseline DC amplitude, Block 526. Both the baseline AC pulse amplitudeand the baseline DC amplitude are determined and stored, Block 528, justshortly following implant of the device using the method of Blocks512-522 as described above. In particular, according an embodiment ofthe invention, the gain coefficient M is calculated for each wavelength,i.e., each light source, Block 524, using the ratio of the baseline ACpulse amplitude and the current average AC pulse amplitude previouslydetermined for each of the light sources in block 522. The offsetcoefficient B is calculated for each wavelength using the differencebetween the product of the current determined gain coefficient M and thecurrent average DC amplitude determined in Block 522 and the baseline DCamplitude. Thus the determination of the gain coefficient M, Block 524,and the offset coefficient B, Block 526, can be made using Equations 2and 3, respectively, set forth below:

$\begin{matrix}{M = \frac{{Baseline}\mspace{14mu} {AC}\mspace{14mu} {Pulse}\mspace{14mu} {Amp}}{{Present}\mspace{14mu} {Ave}\mspace{14mu} {AC}\mspace{14mu} {Pulse}\mspace{14mu} {Amp}}} & {{Equation}\mspace{20mu} 2} \\{B = {{M*\left( {{Present}\mspace{14mu} {Ave}\mspace{14mu} {DC}\mspace{14mu} {Amp}} \right)} - {{Baseline}\mspace{14mu} {DC}\mspace{14mu} {Amp}}}} & {{Equation}\mspace{20mu} 3}\end{matrix}$

Once the gain coefficient M and the offset coefficient B have beendetermined separately for each of the light sources, the device utilizesthe normalization coefficients M and B to correct the respective lightmeasurements subsequently made for each of the light sources for tissueencapsulation, Blocks 420 and 406 of FIG. 4. In particular, for example,the correction for each of the light sources is determined usingEquation 4 set forth below:

X′=M*X−B  Equation 4

Where X is the raw measurement, X′ is the corrected measurement, and Mand B are the current determined correction factors from Blocks 524 and526.

FIG. 4G is a flowchart of a method for correcting sensing by a medicaldevice for the effects of tissue encapsulation according to anembodiment of the invention. It may be desirable to correct the sensingof an optical signal while the patient is physically active, such whenthe patient is walking, for example. As illustrated in FIG. 4G,according to one embodiment, once the device determines it is time toperform an updated calculation of normalization coefficients, Yes inBlock 536, that will be utilized to generate the tissue encapsulationfactor (Block 420 of FIG. 4), a determination is made as to whether thepatient is currently physically active, Block 538, such as whether thepatient is walking, using a vibration sensor, such as an accelerometer,for example. The determination in Block 536 of whether an updatedcalculation of the normalization coefficients is required would besimilar as described above in Block 506 of FIG. 4E. If the patient isdetermined not to be in the desired physical mode, the devices waits apredetermined time period, Block 540, such as one hour for example, andagain makes the determination as to whether the patient is active, Block538. Once the patient is determined to be in the desired physical mode,Yes in Block 538, the device determines the resulting light detected atthe detector to generate a light measurement for each light source,Block 542, which are then stored, Block 544. It is understood that thecalculation of the normalization coefficients is made separately foreach light source, so that if three light sources are utilized, forexample, the calculation is made for each of the three light sources togenerate respective normalization coefficients utilized to correct therespective light measurements subsequently made for each of the lightsources for tissue encapsulation, Blocks 420 and 406 of FIG. 4.

The measurements are made for each light source for a predetermined timeperiod, Block 546, such as 15 seconds for example, or over apredetermined number of steps being taken by the patient, such as 15steps for example. Once light measurements are obtained for the desiredpredetermined time period or over the predetermined number of steps, Yesin Block 546, the device determines a current DC amplitude for each ofthe light sources as being the mean value of the current lightmeasurements during the sampling period, Block 548. A standard deviationor percentile of the current light measurements is determined for eachlight source and a variation between the determined standard deviationsis utilized to generate an AC vibration amplitude for each of thewavelengths, i.e., light sources, Block 550. In another embodiment, theAC values may be determined for each light source in Blocks 548 and 550using the maximum and minimum pulse amplitude associated with the lightmeasurements, during the measurement period as described above inreference to Blocks 522 and 524 of FIG. 4E.

Once the current DC amplitudes and the current AC vibration amplitudeshave been determined for each light source, Blocks 548 and 550,respectively, the device determines a gain coefficient for each of thelight sources, Block 552, using the AC vibration amplitude determined inBlock 550 and a determined baseline AC vibration amplitude, Block 556.In addition, the device determines an offset coefficient, Block 554 foreach of the light sources, using the DC amplitude determined in Block548 and a baseline DC amplitude, Block 556. Both the AC and DC baselinepulse amplitudes are determined and stored, Block 556, just shortlyfollowing implant of the device using the method of Blocks 542-550 asdescribed above. In particular, according an embodiment of theinvention, the gain coefficient M is calculated for each wavelength,i.e., each light source, Block 552, using the ratio of the baseline ACpulse amplitude, Block 556, and the current AC value of vibrationdetermined for each of the light sources in block 550. The offsetcoefficient B is calculated for each wavelength using the differencebetween the product of the current determined gain coefficient M and thecurrent DC amplitude determined in Block 548 and the baseline DCamplitude. Thus the determination of the gain coefficient M, Block 552,and the offset coefficient B, Block 554 are made using Equations 2 and3, respectively, set forth above.

Once the gain coefficient M and the offset coefficient B have beendetermined separately for each of the light sources, the device utilizesthe normalization coefficients M and B to correct the respective lightmeasurements subsequently made for each of the light sources for tissueencapsulation, Blocks 420 and 406 of FIG. 4, using Equation 4 describedabove.

Returning to FIG. 4, the device also corrects the measurements from thelight sources for the effects of mechanical artifacts, Block 408, whichcan include mechanical noise and the effects of mechanical motion andvibration. A variety of different techniques can be used to correct formechanical artifacts. In one technique the measurements of the thirdlight source can be used to compensate for the effects of mechanicalvibration and other noise in the measurements of the first and secondlight sources. Specifically, the measurements of the third light sourcecan be used to remove large excursions in the measurements that can becaused by mechanical vibration and noise.

In general, the excursions created by the noise are faster than therelatively slow changes caused by physiologic perfusion changes.Therefore, in one embodiment the information about the excursions thatis contained in the signal from the third light source may be used toeliminate the noise from the first and second light sources using asuitable noise elimination technique. The corrected first and secondlight source can then be used as described above to determine tissueperfusion.

A variety of different techniques can be used to compensate for theeffects of mechanical vibration on the measurements of the first andsecond light sources using the measurement of the third light source. Asone example, because the effects of noise are tractable between thefirst, second and third measurements, a correlation between themeasurements can be obtained by determining the peak to peak excursionsbetween measurements. The excursions can then be used to determine again factor and the measurements of the third light source to subtractout the noise excursions in the other two measurements. This stepmotion-corrects the other two light measurements. One of themotion-corrected measurements corresponds to the isobestic wavelengthlight, and is dependent on blood volume, but is not dependent on tissueoxygen. The other motion-corrected measurements correspond to light atone of the two non-isobestic wavelengths. Of course, this is just oneexample of how the effects of mechanical motion can be corrected.Additional examples will be discussed below with reference to FIGS. 5and 6.

As illustrated in FIG. 4, the device may also calculate a blood volumeindex, Block 410, and an oxygenation index, Block 412 that are utilizedto provide input to a medical device concerning the hemodynamic statusof the patient, according to an embodiment of the invention. This inputmay be used, for example, to improve the sensing capability of adefibrillator. In general, the blood volume index comprises a metricthat corresponds to the change in the amplitude of the detected opticalsignal from an optical source due to the change in blood volume in thetissue from a baseline volume. Similarly, the oxygenation indexcomprises a metric that corresponds to the change in the amplitude ofthe detected signal from an optical source due to the change in thetissue oxygenation in the tissue from a baseline volume. It should benoted that the volume index and the oxygen index change as a function ofthe wavelength of light emitted from the optical source.

There are a number of ways that the blood volume index and theoxygenation index can be defined. The changes in the amplitude of thelight at a specified wavelength that is measured at the detector areaffected by the change in the blood volume in the tissue and the changein the percentage of that blood that is oxygenated. One embodiment is torepresent the change in the amplitude of the light at a specifiedwavelength that is measured at the detector as a sum of two numbers. Onenumber called the volume index, V_(ind), is dependent on the change inthe volume of blood in the illuminated tissue but not in the change inthe percentage of blood that is oxygenated. The second number called theoxygenation index, Ox_(ind), is dependent on the change in thepercentage of the blood that is oxygenated but not on the change in theblood volume. For example, for an optical source emitting at the redwavelength, this relationship can be illustrated as

I _(red) /I _(bred)−1=(V _(ind))_(red)+(Ox _(ind))_(red)  Equation 5

Where I_(red) is the amplitude of the red light measured at the opticaldetector, I_(bred) is the baseline amplitude of the red light measuredat the optical detector. Without further information, however, it is notpossible to independently calculate (V_(ind))_(red) and(Ox_(ind))_(red).

FIG. 4H is a flowchart of a method for generating a blood volume indexfor each source wavelength in a medical device according to anembodiment of the invention. Once a predetermined triggering event hasoccurred, Block 560, such as a ventricular fibrillation event beingdetected by the device, the device determines the resulting lightdetected at the detector to generate a light measurement for each lightsource, Block 562, which are then stored, Block 564.

The system is set to measure enough samples or over a specified periodso that an accurate decision can be made from the data. For example, ifthe system is supposed to determine whether the patient has lostperfusion due to the onset of fibrillation, the system may be set up sothat the light measurements are taken for a period of ten seconds.

Once the desired number of samples have been obtained, No in Block 566,the device may correct the samples for ambient light, tissueencapsulation and motion, Block 568, as necessary using the techniquesdescribed herein. The decision to employ each of the correction stepsmay be determined by the level of error that each error source causes inthe light measurements. For example, if the system is intended to decideif the volume index has changed by more than 1% in a given time period,then any error source that produces an error in the light measurementsof less than, say, 0.1% can be ignored and the correction step can beskipped. A blood volume index for the isobestic wavelength at time t,[V_(ind)(t)]_(iso) is then determined, Block 570, based on an isobesticbaseline I_(biso) for each of the light measurements generated in Block562 that correspond to the light source associated with the isobesticwavelength I_(iso)(t), using Equation 6 set forth below:

[V _(ind)(t)]_(iso) =[I _(iso)(t)/I _(biso)]−1  Equation 6

The isobestic baseline I_(biso) may be based upon a corrected value ofthe isobestic wavelength taken at a controlled time, such as shortlyafter implant of the device, for example, or the isobestic baselineI_(biso) may correspond to the first light measurement from the lightsource associated with the isobestic wavelength in the series ofmeasurements from Block 562. i.e., the first of the measurements in thecurrent sample period, Block 566. For small changes in blood volume, theblood volume index [V_(ind)(t)]_(iso) is linearly dependent upon thechange in blood volume. The interpretation of the blood volume index isdependent on the technique used to measure the amount of light thatreaches the detector from the light source. If, for example, themeasurement I_(iso)(t) is a direct measure of the amount of lightreaching the detector, then [V_(ind)(t)]_(iso) decreases with increasingblood volume and increases with decreasing blood volume. One example isthat, if the value of the volume index goes negative by more than 0.02,then it may be assumed that that patient has lost perfusion due tofibrillation and defibrillation is necessary.

In one disclosed embodiment, the volume index for wavelengths other thanthe isobestic wavelength are determined by multiplying the isobesticvolume index by a volume correction or scaling factor, C. The scalingfactor C is a predictable value that is defined value that is dependenton the optical properties of blood, the distance from the optical sourceto the optical detector, and the wavelength of the optical source. Forexample, for a red optical source, the volume index isC_(red)×(V_(ind))_(iso). The scaling factor is determined, Block 569,prior to the triggering event, Block 560.

The scaling factor for red light C_(red), can be obtained by the devicein a number of different ways. According to one embodiment, for example,the system can be modeled and C_(red) is then calculated from the modeland stored in the device. According to one embodiment, for example, thevolume correction factor C_(red) is approximately 0.5 for arterial bloodoxygenation of 98% and red and isobestic wavelengths of 660 nm and 800nm, respectively.

According to one embodiment the method to calculate the scaling factorfor red light C_(red) involves applying a least mean squares approach onsignals that regularly change the total blood volume but not thepercentage of blood oxygenation in the tissue that is illuminated by theoptical sensor, as described below. Two such conditions are respirationand arterial pulses. In both cases, the change in blood volume willcause changes in all of the detected signals that are highly correlatedwith each other. In this example, respiration is used to generate thechange in blood volume. Measurement signals can be obtained from eachoptical channel during rest to include a few respiration cycles, such as5-10 seconds long. For example, assume that a two light source systemusing a red light source and an isobestic light source, with thereceived measurement signals being I_(red) and I_(iso), N data pointseach.

A sum of the data points S is defined as:

$\begin{matrix}{S = {\sum\limits_{n = 1}^{N}\left\lbrack {{I_{red}(n)} - {C\left( {{I_{iso}(n)} - {\frac{1}{N}{\sum\limits_{k = 1}^{N}{I_{iso}(k)}}}} \right)}} \right\rbrack^{2}}} & {{Equation}\mspace{20mu} 7}\end{matrix}$

where C is the optimal value of the gain constant. The last sum in theequation can be recognized as the average of I_(iso) signal. Denotingthe average by <I_(iso)> and solving for minimum value of S gives anoptimal value of the scaling factor C_(red) defined as:

$\begin{matrix}{C_{red} = \frac{{{\sum\limits_{n = 1}^{N}\left\lbrack {{I_{red}(n)} \cdot {I_{iso}(n)}} \right\rbrack} -} < I_{iso} > {\sum\limits_{n = 1}^{N}{I_{red}(n)}}}{{{\sum\limits_{n = 1}^{N}{I_{iso}^{2}(n)}} - 2} < I_{iso} > {{\sum\limits_{n = 1}^{N}I_{iso}} + N} < I_{iso} >^{2}}} & {{Equation}\mspace{20mu} 8}\end{matrix}$

The volume index at each source wavelength is calculated according toBlock 570 of FIG. 4H by multiplying the isobestic volume index by thescaling factor for the non-isobestic wavelength.

As illustrated in FIG. 4, the device may also calculate an oxygenationindex, Block 412, that may be utilized by a medical device to decidewhether or not to apply therapy according to an embodiment of theinvention. In general, the tissue oxygenation index is a metric thatrepresents the change from a baseline of the level to which the totalblood supply in the tissue probed by the perfusion sensor is oxygenated.The tissue oxygenation index provides a value that can be used to trackchanges in tissue oxygenation.

FIG. 4I is a flowchart of a method for generating an oxygenation indexin a medical device according to an embodiment of the invention. Asillustrated in FIG. 4I, once a predetermined triggering event hasoccurred, Block 572, such as a possible ventricular fibrillation eventbeing detected by the device, the device determines the resulting lightdetected at the detector to generate light measurements for each lightsource, Block 574, which are then stored, Block 576.

Once the desired number of samples has been obtained, No in Block 578,the device may correct the samples for ambient light, tissueencapsulation and motion, Block 580, as necessary using the techniquesdescribed above. The decision to employ each of the correction steps maybe determined by the level of error that each error source causes in thelight measurements. For example, if the system is intended to decide ifthe oxygen index has changed by more than 1% in a given time period,then any error source that produces an error in the light measurementsof less than, say, 0.1% can be ignored and the correction step can beskipped. An isobestic blood volume index [V_(ind)(t)]_(iso) is thendetermined, Block 582, based on an isobestic baseline I_(biso) and thelight measurements generated from the light source associated with theisobestic wavelength I_(iso)(t) in Block 574, and stored, Block 576. Forexample, using Equation 6 described above, the isobestic blood volumeindex [V_(ind)(t)]_(iso) is the normalized value of the isobesticwavelength I_(iso)(t), generated by the ratio of the isobesticwavelength I_(iso)(t) to the isobestic baseline I_(biso) minus 1(I_(iso)(t)/I_(biso)−1), with the isobestic baseline I_(biso) beingdetermined as described above.

One embodiment for calculating the tissue oxygenation index is tomultiply the isobestic blood volume index by a scaling factor andsubtract the product from the normalized value of the detected signalthat is not at the isobestic wavelength, such as the red signal. Theresulting value would be dependent on changes in tissue oxygen but noton volume. According to FIG. 4I, the volume scaling factor C_(red) isdetermined for the light source associated with the red light wavelengthin Block 583 using one of the techniques described above. The volumecorrection factor C_(red) is determined prior to the Trigger Event,Block 572.

The volume index for a non-isobestic wavelength is generated by, forexample, multiplying the scaling factor for that wavelength times theisobestic volume index, Block 584.

An oxygen index Ox_(ind) is then calculated, Block 586, for each of themeasurements associated with the light source having the red lightwavelength I_(red)(n) based on the correction factor C_(red), a baselinevalue associated with the red light measurements I_(bred), and thecalculated isobestic blood volume index [V_(ind)(n)]_(iso), usingEquation 9 set forth below:

(Ox _(ind))_(red)=[(I _(red)(n)/I _(bred))−1]−[V _(ind)(n)]_(iso) C_(red)  Equation 9

As can be seen from Equation 9, the oxygen index for the red lightmeasurements is determine by subtracting the product of the isobesticblood volume index and the scaling factor for red light([V_(ind)(n)]_(iso)C_(red)) from the normalized value of the detectedsignal for the detected red signal ([(I_(red)(n)/I_(bred))−1).

The baseline value associated with the red light measurements I_(bred)may be based upon a corrected value of the red wavelength taken at acontrolled time, such as shortly after implant of the device, forexample, or the baseline value associated with the red lightmeasurements I_(bred) may correspond to the first light measurement fromthe light source associated with the red wavelength in the series ofmeasurements, I_(red)(n=1) from Block 574. An indication that there isno significant change in tissue oxygenation levels occurs when theoxygen index Ox_(index) is determined to be approximately zero, while adecrease in negative number is an indication of decreasing bloodoxygenation. A decrease in the Ox_(index) may be interpreted as either adecrease in the blood flow through the illuminated tissue or an increasein the oxygen consumed by the illuminated tissue.

In embodiments where three wavelengths of light are used, the sametechnique could also be applied to calculate the change in tissue oxygenin the third light level using a second calibration constant obtainedusing the third light signal and the isobestic light signal.

In one variation on this embodiment, the oxygenation index can be aratio-metric combination of two compensated measurements for the firstand the third light source. This technique provides the ability toamplify the compensated measurements and suppress undesired common-modesignals. For example, the compensated measurements of the first lightsource could be divided by the compensated measurements of the thirdlight source. This would result in magnifying the effects of a change inoxygenation. Additionally, this could improve the quality of signal bycanceling out the effects of some types of noise.

Returning to FIG. 4, with the index calculated, the device determineswhether measurements from additional samples are to be used to calculateadditional blood volume indices and oxygenation indices, Block 414.Typically, it is desirable to calculate multiple oxygenation indexesduring a potential medical event to determine the rate of change inoxygenation of the tissue. Thus, if more calculations are warranted, themethod returns to step 402, where new measurements are taken and a newblood volume index and new oxygenation index is calculated using steps402-412. When multiple blood volume and oxygenation indices have beencalculated, the change in tissue perfusion can be calculated in step416. These changes can then be used by the medical device to determinewhat, if any, action should be taken by the medical device.

As one example, the calculated changes in blood volume and tissueoxygenation indices can be compared to threshold values, respectively,to determine if that change is indicative of a lack of perfusion in thetissue. In this example either a change in oxygen or a change in bloodvolume could indicate a change in perfusion. As one specific example,these threshold values would typically have been previously determinedduring clinical studies in humans to define tolerated and non-tolerateddrops in tissue oxygen, blood volume and/or perfusion.

In another example, the system may combine the signals to generate asingle perfusion index. One simple example is that the blood volumeindex would be compared to its threshold value during a specific timeperiod after the onset of the measurement and the tissue oxygen would becompared to its threshold at other time periods. It is also possible togenerate a perfusion index that is a weighted sum of the blood volumeindex and the tissue oxygen index. This weighting may change as afunction of time after the start of the measurements.

When it is determined that the change exceeds the threshold, the medicaldevice determines what action to take. Again, in one specific examplethe change can be used to determine whether or not defibrillation isdesirable. Specifically, when an ECG based algorithm suspects a VF oruntolerated VT episode, the change in perfusion calculated by method 400can be used to determine if the patient is experiencing a tolerated VTepisode where defibrillation would not be desirable. Of course, this isjust one example of how a medical device, including external andimplantable medical devices, can use a determination of change in tissueperfusion.

As described above, one issue in determining tissue perfusion is thelarge fluctuations in measured light signals that result from mechanicalartifacts. These mechanical artifacts can be caused by the motion ofnearby muscles, or by patient motion causing vibration in thesurrounding tissue. FIG. 5 shows two graphs of measured light signalstaken over 45 seconds for a subcutaneously implanted device in which thetest subject was subjected to manual muscle motion. In these graphs, thelight measurements are in the form of time intervals needed for lightreceived by the detectors to reach a threshold level. The time intervalis the time that light is emitted until the total number of photonsreaching the detector reaches the threshold. Thus, a longer interval isindicative of less transmission through the tissue.

The top graph 502 shows the measured red interval (i.e., from a 660 nmlight source). The bottom graph 504 shows the corresponding measurementsfor an infrared (IR) interval (i.e., from an 880 nm light source). Usingtime intervals as measurements the red and IR intervals are inverselyproportional to tissue transmission at the particular wavelength.Specifically, the sensor turns on the red LED and keeps it on until thetotal number of photons collected by the photo detector reaches athreshold. The time interval that the red LED is ON is called the redinterval. Then the red LED is turned off and IR LED is turned on tomeasure IR interval. This process is continued cyclically.

As can be seen in FIG. 5, mechanical motion causes large fluctuations inboth time intervals. In the illustrated example, the fluctuations in redand IR intervals due to motion are 16 and 17% of mean intervals,respectively. However, the two signals are well correlated.

A variety of different techniques can be used to combine the lightmeasurements to eliminate the effects of mechanical vibration. Any ofthese techniques can be used to implement step 408 of method 400described above. As a first example, the optical signals can beprocessed with a low-pass filter that smoothes out the variations. Thisapproach, however, can cause significant delay in the time to processthe signals and then make a decision.

Other techniques make use of the fact that the variations in the opticalsignals caused by the mechanical motion are very well correlated witheach other. One example is very similar to that already described forseparating the effects of blood volume from tissue oxygen content bycalculating a scaling factor and subtracting a scaled optical signalfrom the others. In this example, the IR signal is used to eliminate themechanical noise from the Red signal. First, a scaling factor iscalculated using the two optical signals. During a short time interval,e.g. one second, the maximum and minimum Red value (I_(Red) _(—) _(Ma)and I_(Red) _(—) _(Min), respectively) and the maximum and minimum IRvalues (I_(IR) _(—) _(Ma) and I_(IR) _(—) _(Min), respectively) aremeasured and the average value of the IR signal, <I_(IR)> is calculated.The scaling factor, C, is defined as:

$C = \frac{t_{IR\_ Max} - T_{IR\_ Min}}{t_{RED\_ Max} - T_{RED\_ Min}}$

The motion compensated red signal is then calculated as:

t _(red) _(—) _(comp) =t _(red)−C·(t _(IR) −<t _(IR)>)

By recalculating the scaling factor and average value during each shortinterval, the information on slow changes in blood volume is maintainedin the compensated signal.

Turning now to FIG. 6, the results of compensation are illustrated ingraph 602. Specifically, graph 602 shows both an uncompensated redinterval in the presence of muscle motion 604 and a red interval afterhaving been compensated for using a scaled IR interval 606. In thisexample, the compensation was performed by multiplying the infraredintervals by a gain constant of 0.256 and subtracting the scaled IRintervals from the red intervals. As can be seen, the effects of motionare significantly reduced by the procedure. Thus, the compensated redintervals 606 are now substantially independent of the effects ofmechanical artifacts. Similarly, the IR measurements can be used toremove mechanical artifacts from a third wavelength, such as from anisobestic wavelength at 805 nm.

Sophisticated algorithms such as adaptive noise cancellation techniquescan also be employed to eliminate noise due to motion from one channelusing another channel that contains correlated noise source due tomotion as well. These techniques lend themselves better to real-timenoise cancellation. This may be advantageous if the correlation of thenoise signals between two channels is slowly changing. However, theycome at a price of increasing signal processing cost and powerconsumption.

In order to compensate for motion and to be able to calculate a bloodvolume index and an oxygenation index, three wavelengths are desirable,one isobestic and two non-isobestic. One of the non-isobesticwavelengths is used to compensate the isobestic and the othernon-isobestic for motion and mechanical vibration. The compensatednon-isobestic wavelength and the isobestic wavelength are then processedto generate the blood volume index and the tissue oxygenation index.

Turning now to FIG. 7, a graph 702 illustrates uncompensated IR timeintervals taken using light at 880 nm. As illustrated in graph 702, thetime intervals have variations due to arterial pulses and variations dueto respiration. Specifically, the smaller higher frequency signals aredue to arterial perfusion. The larger low frequency variations are dueto respiration. After fibrillation, the perfusion stops, leaving thesignal variations due to breathing. Additionally as illustrated, afterfibrillation the lack of perfusion causes the overall, DC level ofoxygenation in the blood to decline.

The variations due to respiration illustrated in graph 702 can be usedto calculate the appropriate gain constant, which will then be used tocompensate for the effects of muscle motion. Specifically, because thevariations due to respiration are similar in effect to the variationsdue to muscle motion, the gain constant calculated to reduce the effectsof respiration caused variation will also reduce the effects of musclecaused variation. Thus, when calculated, the gain constant can be usedto compensate for the effects of muscle motion in the measurements.Additionally, by periodically performing this calculation, the effectsof tissue encapsulation on the sensor can also be corrected for.

For example, the measurement obtained from the reflection at 660 nmwavelength can be compensated via the measurement obtained at 805 nmwhich has relatively low sensitivity to tissue oxygen change. Using thethird light source, the signal could also be compensated by anotherwavelength farther way, preferably at the opposite side of isobesticwavelength (810 nm) such that during an event that causes tissueoxygenation to drop, signal change in primary channel (such as red) canbe maximized. During an event, this calibration gain constant can thenbe used to minimize motion artifacts or can be fed into an adaptivenoise cancellation algorithm as a seed to speed up convergence of thereal-time adaptive algorithm.

As discussed above, one issue in an implanted medical device is changesthat occur to the implant pocket over time. Specifically, due tohealing, fibrous capsule formation and maturation of the area around thedevice, the optical properties of the tissue nearby the perfusion sensorwill change over time. Given that most implantable devices are designedto last for many years it is desirable to compensate for these changes.One suitable technique is to use a periodic calibration of the perfusionsensor that is fast enough to account for changes in nearby tissueoptical properties. For example, by periodically updating the gainconstant using the techniques described above.

As discussed above, in one particular application the perfusion sensoris included in a subcutaneously implanted cardiac device. Themeasurement of perfusion is used to help distinguish between a toleratedventricular tachyarrhythmia (VT) and an untolerated VT. In general, a VTevent is where the heart beats are much faster than normal sinus rhythmand the origin of the rapid beats are within ventricles. An untoleratedVT will cause syncope of the patient due to inadequate perfusion of thebrain. In an untolerated VT, it is desirable to shock the patient toalleviate the tachyarrhythmia, while in a tolerated VT this is generallynot desirable or necessary. Since ECG signal does not tell anythingabout mechanical performance of the heart as a pump, it is difficult forthe implanted device to distinguish between these types of events basedsolely on ECG signals. However, a tolerated VT typically does not resultin a substantial drop in perfusion, while an untolerated VT does. Thus,the perfusion sensor can provide information to the implanted cardiacdevice to distinguish between VT events.

Turning now to FIG. 8, a graph 802 illustrates the change in the redamplitude taken using light at 660 nm. The signal is normalized asI_(red)(t)/Ib_(red)−1. Similarly, graph 804 illustrates the change inthe isobestic amplitude taken using light at 800 nm. As illustrated ingraph 802 and 804, cardiac fibrillation is started at approximately 35seconds, 706, and ends at approximately 65 seconds, 808. Afterfibrillation, the lack of perfusion causes the normalized red amplitude,802, and the normalized isobestic amplitude to change.

Turning now to FIG. 9, the graph 810 shows the isobestic volume indexthat is calculated from the isobestic amplitude using equation 6. Itshould be noted that, for this example, the measurements are defined insuch a way that a decrease in the isobestic volume index indicates adecrease in the blood volume of the sampled tissue. The red oxygenationindex 812 is derived by subtracting the Isobestic Volume Index, which isscaled by a factor C, from the normalized red amplitude 802 from FIG. 8.This procedure was defined by Equation 9. It should be noted that, forthis example, the measurements are defined in such a way that anincrease in the red oxygen index indicates a decrease in the bloodoxygenation of the sampled tissue.

One possible implementation of the perfusion sensor is with afibrillation or untolerated VT detection algorithm that looks for asteady change in the blood volume index or the tissue oxygen index,indicating a severe drop in tissue perfusion. The rate of change couldthen be used to call an event fibrillation or untolerated VT. This couldbe determined experimentally since oxygen is supposed to drop muchfaster with lack of perfusion compared to events such as physicalexercise. One can also employ an automatic algorithm to check for rateof change of compensated red and IR signals during physical exerciseusing activity sensor and rate of change can be determined duringexercise. With that information, a shock threshold can be maintainedabove the rate of change with an adequate safety margin.

Thus, the invention provides a sensor system and method for monitoringchanges in local tissue oxygenation and blood volume, and calculatingthe resulting change in tissue perfusion that is adaptable for use inimplantable medical devices. The tissue perfusion sensor system andmethod provides the ability to determine if the flow of oxygen-richblood through nearby tissue is being maintained, and thus can be used toevaluate the health of the patient's cardiac system. Specifically, bydetermining if perfusion is being maintained the system and method canbe used, along with other sensor measurements, to determine what action,if any, the implantable medical device should take.

While at least one exemplary embodiment has been presented in theforegoing detailed description, it should be appreciated that a vastnumber of variations exist. It should also be appreciated that theexemplary embodiment or exemplary embodiments are only examples, and arenot intended to limit the scope, applicability, or configuration of theinvention in any way. Rather, the foregoing detailed description willprovide those skilled in the art with a convenient road map forimplementing the exemplary embodiment or exemplary embodiments. Itshould be understood that various changes can be made in the functionand arrangement of elements without departing from the scope of theinvention as set forth in the appended claims and the legal equivalentsthereof.

1. A medical device for sensing cardiac events, comprising: a pluralityof light sources capable of emitting light at a plurality ofwavelengths; a detector to detect the emitted light; and a processor togenerate an ambient light measurement in response to ambient lightdetected by the detector, generate a plurality of light measurements inresponse to the emitted light detected by the detector, and adjust theplurality of light measurements in response to the ambient lightmeasurement.
 2. The device of claim 1, wherein the processor generates afirst ambient light measurement prior to the generating of the pluralityof light measurements and a second ambient light measurement subsequentto the generating of the plurality of light measurements.
 3. The deviceof claim 1, wherein the processor generates a first ambient lightmeasurement prior to a first light measurement of the plurality of lightmeasurements, a second ambient light measurement subsequent to the firstlight measurement and prior to a second light measurement of theplurality of light measurements, and a third ambient light measurementsubsequent to the second light measurement.
 4. The device of claim 1,wherein the processor adjusts the generated plurality of lightmeasurements in response to the relative timing of the plurality oflight measurements to a first ambient light measurement and a secondlight measurement.
 5. The device of claim 1, wherein the adjustingcomprises subtracting an average of a first ambient light measurementand a second ambient light measurement from one or more of the pluralityof light measurements.
 6. The device of claim 1, wherein the processoradjusts a light source of the plurality of light sources in response toone of a first ambient light measurement and a second ambient lightmeasurement being greater than the ambient threshold, and, in responseto the one of the first ambient light measurement and the second ambientlight measurement not being greater than the ambient threshold, storesthe first ambient light measurement, the second ambient lightmeasurement, and an intensity of light emitted by the light source ofthe plurality of light sources detected at the detector.
 7. The deviceof claim 6, wherein the adjusting the light source of the plurality oflight sources comprises increasing an amplitude and decreasing a pulsewidth of a light signal emitted by the light source of the plurality oflight sources.
 8. The device of claim 1, wherein the processor adjuststhe detector in response to a first ambient light measurement not beingless than an ambient threshold, and adjusts a first light source of theplurality of light sources in response to a first light measurement ofthe plurality of light measurements being greater than an upper limitthreshold.
 9. The device of claim 8, wherein the processor adjusts thefirst light source in response to an output corresponding to a secondlight measurement of the plurality of light measurements not being lessthan a lower limit threshold, the lower limit threshold being less thanthe upper limit threshold.
 10. The device of claim 9, wherein theprocessor adjusts an output of the first light source in response to theadjusting of the detector and the first light source.
 11. A method ofadjusting sensing in a medical device, comprising: determining ambientlight received at a detector to generate a first ambient lightmeasurement; determining light emitted by a plurality of light sourcesand received at the detector to generate a plurality of lightmeasurements at a plurality of wavelengths; determining ambient lightreceived at the detector to generate a second ambient light measurement;and adjusting the generated plurality of light measurements in responseto the first ambient light measurement and the second ambient lightmeasurement.
 12. The method of claim 11, wherein the first ambient lightmeasurement is generated prior to the generating of the plurality oflight measurements and the second ambient light measurement is generatedsubsequent to the generating of the plurality of light measurements. 13.The method of claim 11, further comprising determining ambient lightreceived at the detector to generate a third ambient light measurement,wherein the first ambient light measurement is generated prior to afirst light measurement of the plurality of light measurements, thesecond ambient light measurement is generated subsequent to the firstlight measurement and prior to a second light measurement of theplurality of light measurements, and the third ambient light measurementis generated subsequent to the second light measurement.
 14. The methodof claim 11, wherein the generated plurality of light measurements areadjusted in response to the relative timing of the plurality of lightmeasurements to the first ambient light measurement and the second lightmeasurement.
 15. The method of claim 11, wherein the adjusting comprisessubtracting an average of the first ambient light measurement and thesecond ambient light measurement from one or more of the plurality oflight measurements.
 16. The method of claim 11, further comprising:determining whether one of the first ambient light measurement and thesecond ambient light measurement is greater than an ambient threshold;adjusting a light source of the plurality of light sources in responseto one of the first ambient light measurement and the second ambientlight measurement being greater than the ambient threshold; and storingthe first ambient light measurement, the second ambient lightmeasurement, and the determined light emitted by the light source of theplurality of light sources in response to the one of the first ambientlight measurement and the second ambient light measurement not beinggreater than the ambient threshold.
 17. The method of claim 16, whereinthe adjusting the light source of the plurality of light sourcescomprises increasing an amplitude and decreasing a pulse width of alight signal emitted by the light source of the plurality of lightsources.
 18. The method of claim 11, further comprising: determiningwhether the first ambient light measurement is less than an ambientthreshold; adjusting the detector in response to the first ambient lightmeasurement not being less than the ambient threshold; determiningwhether an output corresponding to a first light measurement of theplurality of light measurements is greater than an upper limitthreshold; and adjusting a first light source of the plurality of lightsources in response to the first light measurement being greater thanthe upper limit threshold.
 19. The method of claim 18, furthercomprising: determining whether an output corresponding to a secondlight measurement of the plurality of light measurements is less than alower limit threshold, the lower limit threshold being less than theupper limit threshold; and adjusting the first light source in responseto the second light measurement not being less than the lower limitthreshold.
 20. The method of claim 19, further comprising adjusting theoutput of the first light source in response to the adjusting of thedetector and the first light source.